Drug delivery apparatus

ABSTRACT

An implantable drug delivery apparatus for delivering a drug into a bodily fluid in a body cavity of a patient over a period of time includes a variable-volume vessel defining a working chamber for receiving a drug and recirculating a therapeutic fluid. The fluid can contain a bodily fluid, such as, for example, perilymph, and a drug. The device allows for the controlled delivery of the therapeutic fluid to a predetermined location in the bodily cavity of the patient, such as, for example, a cochlea of a human ear.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 11/046,540 filed Jan. 28, 2005, which claims priority to andthe benefit of U.S. provisional patent application Ser. No. 60/540,283,filed Jan. 29, 2004, and U.S. provisional patent application Ser. No.60/602,691, filed Aug. 19, 2004, each application being incorporatedherein by reference in its entirety.

FIELD OF THE INVENTION

The present invention relates generally to the field of drug deliverydevices employing catheters and/or cannulas to transport fluid from areservoir to a patient and, more particularly, to a device forintroducing a drug into patient's bodily fluid, such as, for example,into perilymph in the human ear, as well as to methods for infusingdrugs into cochlea for treatment of hearing loss and other disorders ofhearing and vestibular function.

BACKGROUND OF THE INVENTION

Sensorineural hearing loss (SNHL) is common, and its impact on humancommunication and quality of life, significant. It is estimated thatsome 28 million individuals in the United States suffer from hearingloss. As our population ages, hearing loss prevalence is expected toclimb rapidly, nearly doubling by the year 2030. Causes range fromdegenerative processes associated with aging and genetic disorders toenvironmental exposure to loud sounds and toxic agents. Consequencesrange from moderate communication difficulty and social withdrawal toprofound deafness and its significant challenges. At present, managementof SNHL centers on the use of hearing aids and cochlear implants.However, such treatments cannot address hearing loss prevention, theycannot minimize hearing loss progression and, even with optimal devicefitting, cannot increase a damaged ear's basic capacity. As a result,many users continue to experience significant communicationdifficulties.

Recent advances in the pharmacology and molecular biology of hearinghave revealed new and powerful possibilities for preventing orminimizing hearing loss. The crux of the problem in SNHL is loss of thedelicate cochlear sensory cells that detect the exquisitely smallmechanical vibrations associated with sound. In human ears, once lost ordamaged, these sensory cells do not regenerate and this compromise isoften followed by secondary degeneration of auditory neurons. However,scientists and clinicians are making rapid progress in understanding themolecular mechanisms associated with cochlear and auditory nervedegenerative processes. Additional insight into the molecular signalsinvolved in generating new hair cells is rapidly accumulating, and withthis insight comes the promise of novel and precise drug treatments.Moreover, the extraordinary progress that has been made in defining thegenes involved in a number of human genetic forms of deafness offershope for gene-transfer and molecular approaches to treat these diseases.

For therapies based on these discoveries to become clinically useful, itwill be necessary to develop safe and reliable mechanisms for thedelivery of complex compounds into the inner ear. Direct delivery to thefluids of the inner ear is necessary because of the presence of ablood-labyrinth drug barrier, which is anatomically and functionallysimilar to the blood-brain barrier. That is, through the presence ofso-called ‘tight junctions’ between adjacent cells in the inner ear endorgans, substances outside these organs encounter a substantial physicalbarrier to entry, this protecting the delicate sensory structures withinfrom insult. This ‘protection’, however, also prevents certain moleculeswith potentially therapeutic effect from gaining access to their innerear targets. Prime candidates for exclusion from the cochlea aftersystemic injection are complex molecules, such as proteins and peptides,as well as any molecule that is not lipid-soluble.

Current otologic practice requires drug delivery to the inner ear, butuses inefficient routes. Drugs are commonly delivered systemically, withthe hope that they will find their way to their intended inner eartargets in the form and concentration desired and without serous sideeffects. Systemic corticosteroids, for example, are used in the otologicmanagement of idiopathic sudden and immune-mediated sensorineuralhearing losses. Their clinical usefulness, however, is limited byundesirable side effects arising from the high systemic doses requiredto achieve sufficient cochlear fluid levels of drug to produce theintended inner ear effects.

Local drug application by transtympanic perfusion of the middle ear withthe goal of diffusion through the round window membrane (RWM) into thefluid spaces on the inner ear) was introduced nearly 50 years ago withaminoglycoside treatment of Meniere's disease. This method or somevariant remains in common use in the treatment of inner ear diseases,notably the intractable vertigo that can be associated with Meniere'sdisease, but has been used as well for sudden sensorineural hearingloss, autoimmune inner ear disease, and even tinnitus. Accomplished asan office procedure, a drug is injected through the tympanic membraneinto the middle ear space. The patient then lies with the treated ear‘up’ so that the drug has a better chance of making contact with theRWM, through which the drug must diffuse to gain access to the innerear. With the goal of extending the time of drug availability to theinner ear, newer methods of intratympanic drug delivery have employedseveral strategies to prolong drug contact with the RWM, including,placing absorbent material on or near the TWM and using pump-drivenmicrocatheter systems.

Delivery of drugs to the middle ear reduces systemic side effects, butaccess to the inner ear is unpredictable. Middle ear application hasadvantages over systemic drug delivery, in that drugs so applied canreach their desired targets at higher concentrations and withoutunwanted systemic side effects. The application is straightforward, andcomplications are minimal. A major limitation of these methods, however,is the inability to precisely control the amount of drug that diffusesfrom the middle ear through the RWM into the inner ear. Individualvariation in mucous membrane thickness, mucosal folds and middle earanatomy can have a significant impact on the amount of drug thatultimately enters the inner ear. Some commentators, for example, reportround window niche obstruction in 33% of human ears. This becomes evenmore problematic when considering delivery of coplex macromolecules withlimited diffusion coefficients and those requiring sequenced delivery.Additionally, the bolus application used by certain existing systemsmakes them poorly suited for direct inner ear delivery. Although suchdevices may be useful for delivery of low molecular weight, stable,lipid-soluble compounds like steroids, they would not be suitable forthe delivery of the unstable macromolecules that ultimately will be thetherapeutic compounds with greatest potential benefit.

Direct intracochlear drug delivery, which has been utilized successfullyin animals, has significant potential advantages for therapeuticapplication. The practice of placing drugs of interest within cochlearperilymphatic spaces via a perfusion technique is a method with a longhistory of successful application. When carefully administered, thetechnique itself has been shown to have little effect on a variety ofgross cochlear and neural potentials as recorded from sites within andnear the cochlea. This mode of delivery bypasses the blood-cochleabarrier, allowing drugs to reach their intended targets more directlywith lower doses and fewer non-specific actions. Drugs are largelyunaltered by metabolic changes that inevitably occur with other routesof administration. Drugs perfused into the perilymph compartment ofscala tympani have ready access to the hair cells and synaptic regionsof hair cells, a view supported by investigations in which variousstains demonstrated ready access to structures within the organ of Cortiwhen introduced via the scala tympani perilymph compartment.Additionally, a comparison of the concentrations of cholinergicantagonists required to block the cochlear efferents in vivo and thoseeffective at in vitro isolated outer hair cells shows remarkable closeagreement.

Thus, in order to treat ear disorders, it may often be necessary todeliver therapeutic agents to various ear tissues in a controlled, safe,and efficient manner. For example, a variety of structures have beendeveloped which are capable of delivering/administering therapeuticagents into the external auditory canal of the outer ear. U.S. Pat. No.4,034,759 to Haerr discloses a hollow, cylindrical tube manufactured ofsponge material, e.g. dehydrated cellulose, which is inserted into theexternal auditory canal of a patient. When liquid medicines are placedin contact with the tube, it correspondingly expands against the wallsof the auditory canal. As a result, accidental removal of the tube isprevented. Furthermore, medicine materials absorbed by the tube aremaintained in contact with the walls of the external auditory canal fortreatment purposes.

However, as mentioned above, the delivery of therapeutic agents in acontrolled and effective manner is considerably more difficult withrespect to tissue structures of the inner ear (e.g. those portions ofthe ear surrounded by the otic capsule bone and contained within thetemporal bone which is the most dense bone tissue in the entire humanbody). The same situation exists in connection with tissue materials,which lead into the inner ear (e.g. the round window membrane).Exemplary inner ear tissue structures of primary importance fortreatment purposes include but are not limited to the cochlea, theendolymphatic sac/duct, the vestibular labyrinth, and all of thecompartments (and connecting tubes) that include these components.Access to these and other inner ear tissue regions is typically achievedthrough a variety of structures, including but not limited to the roundwindow membrane, the oval window/stapes footplate, the annular ligament,and the otic capsule/temporal bone, all of which shall be considered“middle-inner ear interface tissue structures” as described in greaterdetail below. Furthermore, as indicated herein, the middle ear shall bedefined as the physiological air-containing tissue zone behind thetympanic membrane (e.g. the ear drum) and ahead of the inner ear.

The inner ear tissues listed above are of minimal size and only readilyaccessible through invasive microsurgical procedures. In order to treatvarious diseases and conditions associated with inner ear tissues, thedelivery of drugs to such structures is often of primary importance.Representative drugs that are typically used to treat inner ear tissuesinclude but are not limited to urea, mannitol, sorbitol, glycerol,lidocaine, xylocaine, epinephrine, immunoglobulins, sodium chloride,steroids, heparin, hyaluronidase, aminoglycoside antibiotics(streptomycin/gentamycin), antioxidants, neurotrophins, nerve growthfactors, various therapeutic peptides, and polysaccharides. Thetreatment of inner ear tissues and/or fluid cavities may involvealtering the pressure, volume, electrical activity, and temperaturecharacteristics thereof. Specifically, a precise balance must bemaintained with respect to the pressure of various fluids within theinner ear and its associated compartments. Imbalances in the pressureand volume levels of such fluids can cause various problems, includingbut not limited to conditions known as endolymphatic hydrops,endolymphatic hypertension, perilymphatic hypertension, perilymphatichydrops, perilymphatic fistula, intracochlear fistula, Meniere'sdisease, tinnitus, vertigo, hearing loss related to hair cell organglion cell damage/malfunction, and ruptures in various membranestructures within the ear.

With respect to existing methods of drug delivery, implantable andexternally mounted drug infusers use a “one-way” infusion system where areservoir empties into the tissue directly or through a catheter. To bepumped along a catheter, however, drugs must have appropriate physicalproperties. For example, it has been determined that dry compounds,which may be more stable than aqueous ones, cannot be used in aconventional infuser. In another example, it has been determined thathighly concentrated compounds may be prohibited because of localreaction at the catheter outlet. Moreover, in the application to innerear diseases, dosage to the relevant tissues of the cochlea can bedifficult or impossible to assess and control by the methods describedabove, and no device has been provided for programmable long-termdelivery, either to the middle ear or inner ear.

Known methods require a relatively complicated mechanism to achievemixing and circulating flow between reservoir and patient. These morecomplicated methods include having two tubes entering the patient,rather than just one, or having a two-way pump, two pumps, or aswitching valve at the pump.

For example, drugs are delivered to the inner ear by infusing the middleear and allowing the medication to diffuse through the local tissue andinto the inner ear. Alternatively drugs are given systemically (e.g.,orally or by injection). For example, U.S. Pat. No. 5,895,372 to Zenner,incorporated by reference herein, discloses an implantable dosagingsystem that injects drugs into the middle ear using a manually operatedpump. As another example, U.S. Pat. No. 6,685,697 to Arenberg et al.,incorporated by reference herein, describes a drug delivery unit forcontrolled delivery of a therapeutic agent to an internal cavity of theear, particularly to the inner ear, that includes carrier media materialcontaining one or more therapeutic agents therein. The carrier mediamaterial is designed to release the therapeutic agents in a controlledmanner over time. The drug delivery unit is shaped and sized forplacement of at least a portion thereof in the round window niche of apatient.

Alternatively, body fluid is caused to circulate through adrug-containing reservoir via a recirculating system having twotubes—one for inflow and one for outflow between reservoir and patient.For example, U.S. Pat. No. 5,643,207 to Rise, incorporated by referenceherein, describes recirculating body fluid through a drug deliverydevice for drug delivery to the brain. As another example, U.S. Pat. No.6,561,997 to Weitzel et al. discloses a circuit for extracorporealtreatment of a body fluid.

As another example, one known perfusion technology involves a cochlearimplant electrode modified to allow intracochlear drug delivery. Inconventional use, the electrode is inserted into the cochlea and used toprovide stimulation to the auditory nerve of severely to profoundlyhearing impaired individuals. The electrode employed for the drugdelivery application, however, contains a removable stylet used forpositioning the electrode during insertion. With the stylet removed, thelumen that remains provides the path for drug delivery. The lumen isconnected to an osmotic or mechanical pump via a connector and shortlength of perfusion tubing.

Notably, existing drug-delivery technology is typically not appropriatefor long-term programmable infusion into the inner ear. The existingapproaches for drug delivery devices include external and implantedinfusers, osmotic pumps, and erodible polymer-drug systems. Thesesystems range from passive devices, which have a low level ofpredictability in their dispense rates, to electronically-controlledrate dispensers and finally to fully programmable infusers. Devicevolumes range from pill size (e.g., those available from OculexPharmaceuticals) to over ten cubic inches, generally depending on theirmaximum dispense volume and sophistication of control. Though small involume, erodible polymer and porous membrane systems (e.g., thoseavailable from IMMED, Inc.) must typically be implemented to deliver aspecific compound, or at best, a set of compounds with similar chemistryand transport properties. They are generally short to medium termdelivery devices (less than six weeks) with unalterable, non-constantdelivery profiles. The existing osmotic pump-based delivery systems(e.g. those available from Alzet International) are similar in terms ofdevice size and lifetime, and they too are capable only of fixed ratedelivery. The various available models trade off device size, lifetime,and delivery rate, depending on the application requirements. Infusertechnology has primarily been developed by Medtronic (Minneapolis,Minn.). Devices such as SynchroMed product offer sophisticated controland are effective for treatment for some disorders such as chronic pain.However, because they use macro scale conventionally fabricated pumps,these systems are relatively large. They are practical only whenimplanted in subcutaneous tissue in the torso.

Emerging microsystems present solution to many previously intractablebioengineering challenges. The extension of micorfabrication methodsfrom integrated circuits to many other applications has spawnedmicroelectromechanical systems (MEMS) devices capable of reproducing thefunctions of conventional sensors and actuators at a fraction of thesize and cost. The resulting miniaturization enables complete systems tobe integrated into devices small enough to be implanted in closeproximity to the organ to be treated. In the case of drug delivery,complex automated dosing regimens can be programmed into the system oreven implemented to respond to sensor input of physiologicalmeasurements. Several technologies have emerged that may allowcontrolled release of drug in dried or lyophilized form from discretecompartments.

And so, as described above, developments in cochlear physiology andmolecular biology allow for new and innovative ways of treating andpreventing sensorineural hearing loss. It is desirable to implement asafe and reliable mechanism for delivering bioactive compounds directlyto the inner ear, e.g., a versatile long-term drug delivery system forthe treatment of inner ear disorders that will have broad applicationand the potential for revolutionizing the treatment of hearing loss.

Thus, it is desirable to provide an implantable long-term drug deliverysystem for treatment of inner ear disorders and prevention ofsensorineural hearing loss, specifically, a versatile device that iscapable of delivering multiple simple and complex molecules over longperiods of time, with capability to control and regulate the sequenceand rate of delivery. Such a device can be useful for treatment ofidiopathic and inflammatory conditions affecting the inner ear,including autoimmune inner ear disease, cisplatinum-induced ototoxicity,and possible Meniere's disease. In addition, a wide spectrum of otherdegenerative inner ear disorders may be amenable to treatment with sucha device, including idiopathic, genetically-based, and age-relatedprogressive sensorineural hearing losses.

SUMMARY OF THE INVENTION

In general, in one aspect, the invention features an implantable drugdelivery apparatus for delivering a drug into a bodily fluid in a bodycavity of a patient over a period of time. The apparatus includes ahollow member that defines at least one lumen for facilitating aunidirectional recirculating flow of a therapeutic fluid through thelumen. The fluid can contain a bodily fluid, such as, for example,perilymph, and a drug. The apparatus also includes a pump, for example asingle unidirectional pump, to control the flow rate of the therapeuticfluid through the hollow member, and an interface member incommunication with at least one lumen of the hollow member. The devicethus allows for the controlled delivery of the therapeutic fluid to apredetermined location in the bodily cavity of the patient, such as, forexample, a cochlea of a human ear.

In various embodiments of the invention, the interface member is influid communication with the bodily fluid in the bodily cavity therebyallowing for drug delivery directly into the bodily fluid. In someversions of these embodiments, the interface member is configured toallow bodily fluid to be periodically drawn from the patient's bodilycavity as the therapeutic fluid is being delivered thereto, allowing fora constant circulation of bodily fluid through the hollow member. Thismay be achieved, for example, by manufacturing the hollow member from aclosed loop of tubing, such as a double-lumen catheter, which travelsfrom the body cavity to the pump. In another embodiment of theinvention, a single-lumen cannula, or other equivalent device, can beused as the interface member to connect the hollow member to the bodycavity. By constructing a network of tubing with selected resistance,tubing compliance, pump rates, and other parameters, a wide range ofdesirable transient and oscillating fluid behavior can be achieved inorder to inject fluid into the internal cavity of the patient andwithdraw fluid from the internal cavity of the patient. In someembodiments of the invention, the fluid capacitance and fluid resistanceof at least one of the hollow member and the interface member areselected to obtain an oscillating flow of the therapeutic fluid throughthe interface member.

One or more implementations of the invention may provide one or more ofthe following features. A drug may be infused directly into the humancochlea. Transport of human perilymph, to and from a drug deliverydevice, may be provided, e.g., for the purpose of delivering therapeuticcompounds to the inner ear. Recirculating fluid may be used to filldepleted volume within the device as a drug is dispensed.

One important feature of the invention is enabling recirculating drugdelivery using a cannula interface to an internal cavity of a patient,such as, for example, cochlea. For example, in one embodiment of theinvention a reciprocating infusion technique can be used to allowrecirculation via a single cannula into a body cavity. Also, among otherbenefits, one of the key features of the invention is providing a devicecapable of reducing net infusion rates without having to reduce the flowrate delivered by the pump. This is particularly important when no pumpis capable of delivering small enough flow rates as required by aphysiological limit. Various embodiments of the invention enablerecirculation and control of very low flow rates (e.g., less than 1microliter/minute) as required in the confined volume of the inner earand other locations.

Delivering fluid to the perilymph temporarily alters its net volume andtherefore results in a change in perilymph pressure. In one embodimentof the invention, the apparatus can be configured to minimize anypotential pressure change within the inner ear throughout infusion of atherapeutic agent. However, it has been discovered that periodicpressure pulses delivered to the inner ear can be therapeutic in thetreatment of inner ear disorders, such as Meniere's disease. Therefore,in certain embodiments of the invention the apparatus can be configuredto manipulate perilymph pressure within the inner ear. This can beachieved by a number of techniques, including, but not limited to,regulating the gauge pressure at the pump inlet or controllably alteringthe volume of fluid delivered at any one time.

In order to deliver a drug to the body cavity, the drug can be dissolvedin the therapeutic fluid, such as a bodily fluid, within therecirculating flow of the hollow member. Dependent upon the requirementsfor treatment of a patient, the concentration of drug dissolved in thetherapeutic fluid may either be varied or held constant. Storage ofadditional drug to be dissolved in the therapeutic fluid can befacilitated through the addition of a reservoir member for storing thedrug, which can include one or more chambers and which is connected tothe hollow chamber recirculating the therapeutic fluid through theapparatus. Further control of the drug infusion process can be achievedby adding additional chambers to the reservoir, allowing for morecontrollable mixing of the drug and therapeutic fluid, and/or the mixingof additional drugs to the therapeutic fluid.

To facilitate the recirculation of the therapeutic fluid through thehollow member a pump can be included within the apparatus. In oneembodiment of the invention this pump may be a microelectromechanical(MEMS) microfluidic pump. The pump can be operated at a predeterminedfrequency, which can be either substantially constant or modulateddepending upon the requirements of the system. In a particularembodiment, a flow rate of less than about one microliter per minute.

In various embodiments, control of the flow pattern of the therapeuticfluid through the hollow member is implemented through the use of acontrol system in electric communication with the pump. Performanceparameters regarding the flow pattern of the therapeutic fluid can alsobe detected by the addition of sensors to the apparatus. The informationfrom these sensors can then be transmitted to a remote device throughthe use of receiving and transmitting electronics in both the remotedevice and the apparatus. This configuration can also allow electricsignals to be sent from the remote device to the apparatus. The additionof a memory element to the control system for the pump can also beadvantageous for such purposes as monitoring the performance of thedevice and storing control information. In one embodiment, a sensor canbe used to detect and, optionally, control flow pattern parameters ofthe therapeutic fluid.

In various embodiments, the apparatus includes a regulating system formaintaining an optimal drug delivery rate. The regulating system is incommunication with at least one the pump and the reservoir. For example,the regulating system may include a sensor for periodically measuringconcentration of drug in the therapeutic fluid and/or the bodily fluidand transmitting the measured value of the concentration to theregulating system. In some embodiments, a biosensor detects a level of aparticular molecule of the drug and thereby automatically determine thequantity of drug to release from the reservoir. Also a sensor canmeasure the concentration of drug in the bodily fluid, such asperilymph, and provide feedback to regulate the drug release rate.

In a particular embodiment of the invention, the apparatus is configuredto allow for the long term, for example greater than one year, deliveryof a therapeutic fluid to the inner ear of a human. In this embodiment,at least part of the apparatus can be shaped and dimensioned to fitwithin the mastoid cavity of a human patient. In one version of thisembodiment, the interface member to be in fluid communication with acochlea of the ear. It will therefore be possible, in this and otherembodiments, to use human perilymph as the bodily fluid.

In a specific exemplary embodiment, the invention features an implantedapparatus that fits within the mastoid cavity of humans. The apparatuscontains an externally-programmable pump to recirculate perilymph, anintracochlear catheter inserted into the scala tympani through acochleostomy having a cannula in communication with a body fluid of thepatient, a mixing chamber with externally programmable delivery ofconcentrated bioactive compounds, and sensors for detecting andtransmitting flow and pressure information. The ultra-miniaturizedapparatus is a complete, long-term (one year and greater) deliverysystem, containing therapeutic compound, dispensing mechanism, controlelectronics, and power supply.

Alternatively, the apparatus according to the invention can beimplemented to deliver drugs to a bodily cavity, such as the cochlea ofa human ear, without the need for a recirculating fluid flow within thehollow member connecting the pump to the interface member. Generally, inanother aspect, the invention features an implantable drug deliveryapparatus for delivering a drug into a bodily fluid in a bodily cavityof a patient over a period of time, which includes a hollow memberdefining at least one lumen for facilitating a flow of a therapeuticfluid therethrough, the therapeutic fluid containing a first volume ofthe bodily fluid, such as, for example, perilymph, and the drugcontained therein; a pump for controlling a flow rate of the therapeuticfluid through the hollow member; and an interface member in fluidcommunication with the at least one lumen of the hollow member and thebodily fluid in the bodily cavity, such as, for example, a humancochlea, for delivering at least a portion of the therapeutic fluid intothe bodily fluid in a predetermined location in the bodily cavity. Invarious embodiments of the invention the at least one of the pump andthe hollow member is shaped and dimensioned to fit within a mastoidcavity of a human.

In general, in yet another aspect, the invention features a method fordelivering a drug into a bodily fluid in a bodily cavity of a patientover a period of time, which includes the steps of drawing a firstvolume of the bodily fluid from the patient's bodily cavity; mixing thedrug with the bodily fluid thereby obtaining a therapeutic fluid; andreleasing a first volume of the therapeutic fluid into the bodilycavity.

In various embodiments, the method also includes the step of providing ahollow member defining at least one lumen in communication with thebodily cavity, such as, for example, human cochlea and causing aunidirectional flow of the bodily fluid through the hollow member. Themethod may further include the steps of controllably recirculating thetherapeutic fluid through the hollow member, altering concentration ofthe drug in the therapeutic fluid during recirculation thereof throughthe hollow member, and drawing a second volume of the bodily fluid fromthe bodily cavity. In one example embodiment of the invention, the flowproperties of the bodily fluid through the apparatus can depend upon thecapacitance and resistance of the hollow member and interface member.Careful selection of the capacitance and resistance can provide anoscillating flow through a single interface member. In some embodiments,the flow rate of the therapeutic fluid during recirculation through thehollow member is substantially constant. The therapeutic fluid mayinclude a solution of the drug in the bodily fluid, for example, humanperilymph. In a particular embodiment, the method includes controllablyaltering fluid pressure within the human cochlea, for example,increasing fluid pressure within the cochlea by controllably alteringthe first volume of therapeutic fluid.

In still another aspect, the invention is directed to a drug deliveryapparatus for delivering a drug into a bodily fluid in a bodily cavityof a patient over a period of time. The apparatus includes avariable-volume vessel that defines a working chamber for receiving adrug and recirculating a therapeutic fluid. The therapeutic fluid mayinclude a first volume of the bodily fluid and the drug containedtherein. The apparatus further includes an actuator for varying thepressure within the chamber by altering the volume of the chamber; andan interface member in fluid communication with the chamber and thebodily cavity for periodically delivering at least a portion of thetherapeutic fluid to a predetermined location in the bodily cavity anddrawing a second volume of the bodily fluid from the bodily cavity.

Various embodiments of this aspect of the invention include thefollowing features. The interface member can be, or may include, acannula. The drug can be dissolved in the first volume of the bodilyfluid. The concentration of the drug in the therapeutic fluid may varyduring recirculation thereof through the working chamber. The apparatusmay also include a reservoir member for storing the drug, which definesa storage chamber in communication with the working chamber. Thevariable-volume vessel may have a slidably movable wall or a wall havinga deflectable portion, for example, a flexible membrane. The actuator,for example, selected from the group consisting of: electromagnetic,pneumatic, or hydraulic devices, may periodically increase and decreasethe volume of the chamber by at least one of deflecting or slidablymoving at least a portion of at least one wall of the variable-volumevessel. The apparatus may also include a regulating system incommunication at least with the actuator for maintaining a desirabledrug delivery rate and/or a control system in electric communicationwith the actuator for controlling a flow pattern of the therapeuticfluid through the working chamber. The variable-volume vessel and theactuator can be shaped and dimensioned to fit within a mastoid cavity ofa human.

In yet another aspect, the invention features a method for delivering adrug into a bodily fluid in a bodily cavity of a patient over a periodof time. The method includes the step of drawing a first volume of thebodily fluid from the patient's bodily cavity through an interfacemember directly into a variable-volume vessel defining a workingchamber. The method further includes mixing the drug with the bodilyfluid in the working chamber thereby obtaining a therapeutic fluid; andincreasing a pressure within the working chamber to release a firstvolume of the therapeutic fluid into the bodily cavity.

BRIEF DESCRIPTION OF THE DRAWINGS

The objects and features of the invention can be better understood withreference to the drawings described below, and the claims. The drawingsare not necessarily to scale, emphasis instead generally being placedupon illustrating the principles of the invention. In the drawings, likenumerals are used to indicate like parts throughout the various views.

FIG. 1 depicts a sketch of a human inner ear with an implanted drugdelivery system, in accordance with various embodiments of theinvention.

FIG. 2A is a schematic view of an exemplary drug delivery apparatus thatincludes a pump, reservoir, electronics and battery system, inaccordance with one embodiment of the invention.

FIG. 2B depicts a sketch of a exemplary drug delivery apparatusimplanted in the mastoid cavity of a human ear.

FIG. 3A is a schematic view of a recirculating drug delivery apparatusin accordance with some embodiments of the invention.

FIG. 3B-3C depict schematic diagrams for the drug delivery apparatus ofFIG. 1A.

FIG. 4A is a plot of an example pump flow output for a pump operating ata constant frequency, in accordance with one embodiment of theinvention.

FIG. 4B is a plot of an example flow rate for a pump which isperiodically turned on and off at a frequency lower than the pump cyclefrequency, in accordance with one embodiment of the invention.

FIG. 5A is a plot of an example output flow for one example deliverysystem design with a pump operating at a constant frequency, inaccordance with one embodiment of the invention.

FIG. 5B is a plot of a second example output flow for one exampledelivery system design with a pump operating at a constant frequency, inaccordance with one embodiment of the invention.

FIG. 6A is a plot of an example output flow for one example deliverysystem design with a pump which is periodically turned on and off at afrequency lower than the pump cycle frequency, in accordance with oneembodiment of the invention.

FIG. 6B is a plot of a second example output flow for one exampledelivery system design with a pump which is periodically turned on andoff at a frequency lower than the pump cycle frequency, in accordancewith one embodiment of the invention.

FIGS. 7A-7B are together a table of performance data for exemplaryembodiments of the invention.

FIGS. 8A-8B are schematic views of a drug delivery apparatus inaccordance with alternative embodiments of the invention.

DETAILED DESCRIPTION

As discussed above, conventional drug infusers utilize macroscalemachined components to pump liquid drugs from a reservoir. The inventionprovides for replacing these components with a synthesis of micropumpsand MEMS solutions for drug storage and release, which will result insmaller devices with greater functionality. This will enable the openingof the inner ear and other previously inaccessible locations in the bodyfor new direct treatment, without the side effects of systemic delivery.

Microfluidics and microelectromechanical systems (MEMS) capability canbe used for drug delivery applications, to allow or provide a controlledrate, low drug volume, and/or liquid formulation (e.g. for animplantable inner ear delivery system). In an example embodiment, afluidic system having a closed loop microfluidic flow controller can beused with animal test apparatus. In one embodiment of the currentinvention, an implanted recirculating delivery system can be used intherapy for hearing loss and Meniere's disease. An example deliverysystem may employ a number of commercially available pumps, such as, butnot limited to, a Wilson Greatbatch insulin pump or MEMS pump, such asthose available from Debiotech.

In some embodiments, the micromechanical device for intracochlear drugdelivery discloses utilizing a surgical approach that is similar tocochlear implantation, but minimizes cochlear insult. The implementationconcept includes a double lumen intracochlear catheter inserted intoscala tympani through a cochleostomy adjacent to the round window. Inits implanted position, it is similar to cochlear implants that alsotraverse the tympanomastoid cavity with electrodes positioned within thecochlea, except that the depth of insertion is much less.

In accordance with the invention, drug delivery to the ear relies on amethod in which a recirculating stream of fluid from the patient ispassed through a device and is infused remotely rather than within thetissue, which enables recirculation and control of very low flow rates(e.g., less than 1 microliter/minute) as required in the confined volumeof the inner ear. A specific application with respect to inner eardiseases provides for direct infusion of the cochlea through a catheter,using an implanted device to programmably and continually deliver drugsthrough the catheter.

The recirculating fluid permits the reservoir to contain a highlyconcentrated solution, and therefore can potentially produce a devicethat operates for years without refilling. This greatly reduces the riskof microbial contamination during refill. Another benefit is using avehicle that is inherently biochemically compatible. In addition, theperilymph may circulate through the catheter at a rate that isindependent of the drug delivery rate. Thus these parameters can beoptimized separately. It is likely that frequent circulation of theperilymph will maintain patency in the catheter, whereas a slow one-waydrug infusion would occlude. Finally, because there is controlled supplyof liquid solvent, it is not necessary to use a liquid drug reservoir.The drug storage could take any number of forms, such as microchiparrays, bio-erodible polymers, or even hybrid combinations of these drugdelivery methods.

In a specific exemplary embodiment, a microfluidic pump recirculateshuman perilymph, which is withdrawn and returned to the inner earthrough a catheter, implanted through the round window membrane oradjacent tissue. Drugs are injected into this recirculating stream fromone or more reservoirs by one or more microvalves and/or one or moreother drug release methods.

As used herein, the term “drug” is understood to mean any natural orsynthetic, organic or inorganic, physiologically or pharmacologicallyactive substance capable of producing a localized or systemicprophylactic and/or therapeutic effect when administered to an animal. Adrug includes (i) any active drug, (ii) any drug precursor or pro-drugthat may be metabolized within the animal to produce an active drug,(iii) combinations of drugs, (iv) combinations of drug precursors, (v)combinations of a drug with a drug precursor, and (vi) any of theforegoing in combination with a pharmaceutically acceptable carrier,excipient, or formulating agent.

The drug or drugs of interest may be stored in the apparatus either inpure form or as a formulation, for example, in combination with apharmaceutically acceptable carrier or encapsulated within a releasesystem. A release system can include a matrix of a biodegradablematerial or a material which releases incorporated drug by diffusion.The drugs can be homogeneously or heterogeneously distributed within therelease system. A variety of release systems may be useful in thepractice of the invention, however, the choice of the appropriate systemwill depend upon rate of drug release required by a particular drugregime. Both non-degradable and degradable release systems can be used.Suitable release systems include polymers and polymeric matrices,non-polymeric matrices, or inorganic and organic excipients and diluentssuch as, but not limited to, calcium carbonate and sugar. Releasesystems may be natural or synthetic. However, synthetic release systemsare preferred because generally they are more reliable, morereproducible and produce more defined release profiles. The releasesystem material can be selected so that drugs having different molecularweights are released from a particular cavity by diffusion through ordegradation of the material. Biodegradable polymers, bioerodiblehydrogels, and protein delivery systems currently are preferred for drugrelease via diffusion or degradation.

Representative synthetic, biodegradable polymers include, for example:polyamides such as poly(amino acids) and poly(peptides); polyesters suchas poly(lactic acid), poly(glycolic acid), poly(lactic-co-glycolicacid), and poly(caprolactone); poly(anhydrides); polyorthoesters;polycarbonates; and chemical derivatives thereof (substitutions,additions of chemical groups, for example, alkyl, alkylene,hydroxylations, oxidations, and other modifications routinely made bythose skilled in the art), copolymers and mixtures thereof.Representative synthetic, non-degradable polymers include, for example:polyethers such as poly(ethylene oxide), poly(ethylene glycol), andpoly(tetramethylene oxide); vinyl polymers-polyacrylates andpolymethacrylates such as methyl, ethyl, other alkyl, hydroxyethylmethacrylate, acrylic and methacrylic acids, and others such aspoly(vinyl alcohol), poly(vinyl pyrolidone), and poly(vinyl acetate);poly(urethanes); cellulose and its derivatives such as alkyl,hydroxyalkyl, ethers, esters, nitrocellulose, and various celluloseacetates; polysiloxanes; and any chemical derivatives thereof(substitutions, additions of chemical groups, for example, alkyl,alkylene, hydroxylations, oxidations, and other modifications routinelymade by those skilled in the art), copolymers and mixtures thereof.

Preferably, the storage capabilities of the apparatus are such that itholds sufficient amount of the drug to provide a continuous deliveryover the extended delivery period, e.g., several weeks, months, or evenlonger. The storage volume needed thus depends on characteristics suchas drug solubility, drug delivery rate, period of delivery, drug's halflife, etc. Once implanted, the device continuously delivers the drug forprolonged period of time until replenishment.

In various embodiments of the invention, communication with a remotedevice external to the patient's body and capable of controlling of theinfusion rate allows for modification of the therapy in response to apatient's symptoms and reactions. This feature may include control ofthe recirculation rate to allow different dosage schemes, such as, butnot be limited to, either steady low concentrations or intermittent highconcentrations of drugs. Variation of the dosage based on the time ofday can also be desirable.

In addition to performance features, a number of safety features mayalso be included in embodiments of the invention. Example features mayinclude, but not be limited to, automatic shutoff control if pressure orflow sensors give abnormal readings, self-diagnostic routines which mayrun automatically or upon prompting from an external controller. In oneembodiment of the invention, telemetry can enable a physician tointerrogate settings, identify low battery or other alarm signals, andobtain device identification or serial number. A clinician maycommunicate with the device by means of a hand-held module connected toa personal computer, or through another analogous communication device.

The ability to communicate with implanted electronic devices has beenwell established over the last 25 years (e.g. with pacemaker systems).As such, communicating with and controlling the drug delivery devicedoes not pose a major problem. Nonetheless, the communication subsystemmust guarantee reliable and robust operation, since minimal service andadjustment is possible after installation.

As a result of its ubiquitous application, communication via thewireless RF technique offers one approach for remote communication. Inaddition to enabling a small low-cost device, the RF technique alsoprovides a convenient means by which the battery energy may bereplenished. Although recent studies have concentrated on frequenciesabove a few hundred megahertz, these studies have been motivated by theneed to distribute real-time image information. The bandwidthrequirements for the drug delivery device are much more modest. Afrequency of 10 MHz helps minimize attenuation due to skin effect, whileat the same time allows use of a small, low profile antenna.

Several additional physical means are also available for couplingcommunication signals from the implanted device to an externalinterrogator or programmer. In one embodiment of the invention,mechanical (acoustic) waves may provide a communication mechanism. Theacoustic technique is enabled by the recent availability of miniaturetransducers fabricated with MEMS technology. Further embodiments mayinclude, but not be limited to, the use of optical means or directvolume conduction to communicate with an implanted device.

Referring to FIG. 1, in one embodiment, an implanted recirculatingdelivery system directs fluid to and from the cochlea of a human ear 10.A double lumen catheter 12 is implanted a body and is in communicationwith the vestibule 14 and cochlea 16 of the inner ear. This arrangementallows a fluid to recirculate between the cochlea 16 and an external orinternally planted pump (not shown).

An exemplary embodiment of the invention with an electronic deviceimbedded within the mastoid cavity of a human ear can be seen in FIGS.2A-2B. In FIG. 2A, a device 20 includes a micropump 22 connected to areservoir 24. The flow rate produced by the pump 22, and the rate atwhich a drug is released by the reservoir 24, can be controlled bycontrol system 26 integrated within the device 20. Power can be suppliedto the system through a battery 28, which can also be imbedded in thedevice 20. Alternative embodiments of the device 20 may incorporateadditional features, such as but not limited to further reservoirs oradditional electronic features, but can also be simplified by removingattachments shown herein, such as the reservoir 24. For example, drugstorage within the device can be achieved through a number of methodssuch as, but not limited to, the use of a fluid chamber with a valveconnection, the addition of bio-erodible polymers, the addition ofmultiple reservoirs 24 containing multiple drugs, and the addition ofstorage devices capable of delivering solid or powdered drugformulations.

The device 20 shown in FIG. 2A can be seen implanted within the mastoidcavity 30 of a human ear, in accordance with one embodiment of theinvention. In this embodiment, the device 20, incorporating themicropump 22, reservoir 24, control system 26, and battery 28, isimplanted behind the pinna 34 of a human ear, within the mastoid cavity30. The device is connected to a double-lumen catheter 32, whichconnects to an interface member, in this case a cannula 36, which isimplanted into the vestibule 14 of a human ear, thus allowing fluidcommunication with a cochlea 16.

Various configurations of the device allow a drug, or drugs, to be mixedwith the therapeutic fluid recirculating within the double-lumencatheter 32. Depending upon the requirements of the system, the infusionof a drug into the therapeutic fluid can be constant or modulated. Theflow rate of the therapeutic fluid within the system can also becontrolled through the control of the micropump 22, which can either beheld at a substantially constant frequency or modulated. The controlsystem 26 in the device can control the flow and infusion rate, and alsoprovides the possibility of monitoring the performance of the device 20,send information regarding the flow parameters to a remote device, andreceive information from a remote device. In various embodiments, thedevice includes a regulating system that is used to determine optimaldrug delivery rates. In some embodiments, the regulating system is partof the control system 26. In one particular embodiment, a biosensor ofthe regulating system detects a level of a particular molecule of thedrug and thereby enables the regulating system to automaticallydetermine the quantity of the drug to release from the reservoir. Also,a sensor of the regulating system could also measure the concentrationof drug in the perilymph and provide feedback to regulate the drugrelease rate from the reservoir or increase the flow rate by the pump.

A schematic for the basic fluid circuit is shown in FIGS. 3A to 3C.Referring to FIG. 3A, in one embodiment, a drug delivery system 40 hasbeen designed without a distinct supply reservoir. As a result, itrecirculates a constant net volume of fluid through a loop of tubing 42driven by a micropump 44. The recirculating stream communicates througha lumen of a cannula 46 with the cochlea 48, depicted here forsimplicity as an open reservoir containing fluid 50. Delivery occursthrough transport outside of the system: fluid expelled during the firsthalf pump cycle equilibrates with the fluid in the outside reservoir,either through diffusion or mixing, thus the fluid drawn in during thenext half cycle is less concentrated and net delivery occurs, albeitdecreasing over time. In various embodiments, design of the system 40enhances mixing by achieving an oscillatory flow of sufficient amplitudeto completely expel the fluid contained in the cannula 46 during acycle. Otherwise “fresh” compound would not be delivered each cycle, ineffect, mixing would largely be dominated by diffusion in the smallvolume of the cannula 46.

FIGS. 3B and 3C, respectively, depict a plumbing diagram for therecirculating fluidic delivery system and its equivalent lumped-elementelectric circuit schematic. FIG. 3B depicts a schematic diagram for thesystem 40, with the addition of a pressure gauge 52. The pressure gauge52 is connected to the feed leg of the hollow member, which comprisestwo sections of differing diameter 54 and 56. The return leg of thehollow member comprises the two sections of differing diameter 58 and60. The hollow member connects through a T-junction 62 to the cannula46. FIG. 3C depicts a circuit representation of the system of FIG. 3B.Here, the resistance of the sections of each hollow member section 54,56, 58, and 60 are shown, along with the resistance within the cannula46 and the capacitance in the feed and return legs 64 and 66.

By careful selection of the geometric properties of the cannula 46 andhollow member sections 54, 56, 58, and 60, the flow pattern propertieswithin the system, and the resulting drug delivery rates to the cochlea48, can be controlled. In a particular embodiment of the invention,selection of the systems geometric properties and the operationproperties of the micropump 44 can produce a reciprocating flow withinthe system. In this configuration, the fluid capacitance and fluidresistance of within the delivery system can be selected and,optionally, controllably altered, to provide an oscillating flow througha single cannula 46. This flow regime can have a number of importantbenefits, such as, but not limited to, improving mixing of the drug andperilymph within the delivery system and cochlea 48, carefullycontrolling the rate of drug delivery to the cochlea 48, and helping toavoid occlusion within the tubing. This configuration also allows for atransport of fluid into and out of the cochlea 48 using only a singleinterface member.

In some embodiments, the micropump driving the fluid is a reciprocatingsolenoid pump (such as a Wilson Greatbatch WGL 05) with a 0.5 uL fixedstroke volume operating up to 20 psi. The transition time of the pumpstroke is preferably much smaller than the pump cycle time, which is0.33 sec minimum (3 Hz maximum pumping frequency). The nominal feed andreturn tubing between the pump and T-junction are each approximately 50cm long with negligible resistance, having an I.D. of 1.0 mm. Thesetubes may function as the primary source of compliance (CF and CRdescribed below) and could vary in material from silicone (modulus ˜10MPa) to PEEK (modulus 1 GPa). The T-junction capillaries are rigid(fused silica). The tubes represented by RFT and RRT should have I.D.less than 250 um (not necessarily equal) and length of at least 10 mm.The cannula 46 is assumed fixed, because of surgical constraints, withI.D. 75 um and length 20 mm.

To satisfy the above condition, one half of a flow cycle must generate afluid flow volume of at least that of the mixing tube volume.$V_{M} = {{\frac{\pi}{4} \cdot D_{IM}^{2} \cdot L_{M}} = {0.088\quad{µL}}}$

Given the circuit configuration, it is difficult to achieve this withoutsome capacitance in the system. Specifically, with the fluidiccapacitors shown in FIG. 3C removed, there is no loop that includes themixing output leg RM through which fluid can flow. Equivalently, thereis no storage capability in the pump loop which allows fluid to bestored in such a way that the flow rates in the T-feed and T-returnsections can be unequal at the same instant in time, which is the onlycondition under which fluid may flow in the cannula.

In one embodiment of the fluidic delivery system described withreference to FIGS. 3A-3C, the micropump can be configured to operatecontinuously at a predetermined frequency. In a second embodiment, themicropump input can be modulated so that it periodically turns on andoff at some frequency much lower than the pump cycle frequency, and alsomore slowly than the largest system time constant.

In order to analyze the system described in FIGS. 3A to 3C, a number ofsystem parameters must be calculated for the component geometry andproperties, and a number of simplifying approximations must be made. Forexample, the pump pulse time is estimated to be of the order ofmilliseconds. Also, the resistance to fluid flow of a tube with circularand constant cross section can be given by; $\begin{matrix}{R = \frac{128 \cdot \eta \cdot L}{\pi \cdot D_{I}^{4}}} & \left( {{Formula}\quad 2} \right)\end{matrix}$where η is the dynamic viscosity, L the tube length and D_(I) the innerdiameter.

For an expandable piece of tubing, the capacity to store fluid can beapproximated by; $\begin{matrix}{{C \equiv \frac{\mathbb{d}V}{\mathbb{d}P}} = \frac{\pi \cdot L \cdot D_{I}^{3}}{2 \cdot E_{Y} \cdot \left( {D_{O} - D_{I}} \right)}} & \left( {{Formula}\quad 3} \right)\end{matrix}$where E_(Y) is the elastic modulus, D_(O) is the outer diameter, andD_(I) again refers to the tube inner diameter. Alternatively, to use thecompressibility of a length of air bubble in a portion of tubing, thecapacitance can be described approximately by; $\begin{matrix}{C = \frac{L_{0} \cdot \pi \cdot D_{I}^{2} \cdot P_{0}}{4 \cdot P^{2}}} & \left( {{Formula}\quad 4} \right)\end{matrix}$where L₀ is the length of the bubble when at pressure P₀, and P is thebubble pressure. It should be noted that this expression describes anon-linear element (i.e. it is dependent on the pressure). For analysis,the average pressure of the bubble (i.e. P=P_(avg)) gives reasonablyaccurate estimates of the bubble capacity as long as the average islarge compared to its maximum deviation from that average.

Laplace domain analysis of the circuit in FIG. 3A-3C yields the transferfunction; $\begin{matrix}\begin{matrix}{\frac{I_{0}}{I_{S}} = \frac{{- A_{0}} \cdot \omega_{n}^{2} \cdot s}{\left( {s^{2} + {2 \cdot \zeta \cdot \omega_{n} \cdot s} + \omega_{n}^{2}} \right)}} \\{= \frac{{- A_{0}} \cdot \omega_{n}^{2} \cdot s}{\left( {s + \omega_{H}} \right) \cdot \left( {s + \omega_{L}} \right)}}\end{matrix} & \left( {{Formula}\quad 5} \right)\end{matrix}$where I₀ is the fluid flow through the output tube, I_(S) is the sourceflow, and the system gain, undamped natural frequency, damping ratio,and high and low frequency poles are given respectively by;$\begin{matrix}{A_{0} = {{R_{FT} \cdot C_{F}} - {R_{RT} \cdot C_{R}}}} & \left( {{Formula}\quad 6} \right) \\{\omega_{n} = \left\lbrack {C_{F} \cdot C_{R} \cdot \left( {{R_{FT} \cdot R_{RT}} + {R_{FT} \cdot R_{M}} + {R_{M} \cdot R_{RT}}} \right)} \right\rbrack^{- \frac{1}{2}}} & \left( {{Formula}\quad 7} \right) \\{\zeta = \frac{\begin{matrix}{\omega_{n} \cdot} \\\left( {{R_{M} \cdot C_{R}} + {R_{FT} \cdot C_{F}} + {R_{RT} \cdot C_{R}} + {R_{M} \cdot C_{F}}} \right)\end{matrix}}{2}} & \left( {{Formula}\quad 8} \right) \\{\omega_{H} = {{{\left( {\zeta + \sqrt{\zeta^{2} - 1}} \right) \cdot \omega_{n}}\quad\omega_{L}} = {\left( {\zeta - \sqrt{\zeta^{2} - 1}} \right) \cdot \omega_{n}}}} & \left( {{Formula}\quad 9} \right)\end{matrix}$

It can be shown, by taking partial derivates of Formula (8) with respectto the various circuit elements, that the damping ratio ζ for thissystem is always greater than or equal to one, and in fact is only equalto one in two trivial non-useful scenarios, and thus the system neverhas an under-damped, decaying-oscillation response to an impulse or unitstep input.

FIG. 4A depicts a qualitative plot of the pump flow output for theconfiguration wherein the micropump is configured to operatecontinuously at a predetermined frequency. Here, the cycle frequency is1 Hz and the pulse time is 0.05 sec. In this operating mode, the systemdesign time constants are large compared to the pulse time but smallcompared to the pump cycle period. As a result, the input can be modeledas an impulse function. A single pulse of the pump would be expected togenerate a transient flow event such that the total volume exchangeduring that event exceeded that given by the above stated Formula (1).

The volume impulse response is given by; $\begin{matrix}{V_{imp} = {\frac{V_{stroke}{A_{0} \cdot \omega_{n}}}{2 \cdot \sqrt{\zeta^{2} - 1}} \cdot \left( {{\exp\left( {{- \omega_{H}} \cdot t} \right)} - {\exp\left( {{- \omega_{L}} \cdot t} \right)}} \right)}} & \left( {{Formula}\quad 10} \right)\end{matrix}$where, as mentioned above, it is assumed that the stroke volume isdelivered in a time interval small compared to all system timeconstants. This results in a maximum volume exchange of $\begin{matrix}\begin{matrix}{V_{cycI} = {\frac{V_{stroke}{A_{0} \cdot \omega_{n}}}{2 \cdot \sqrt{\zeta^{2} - 1}} \cdot}} \\{\left\lbrack {\left( \frac{\omega_{H}}{\omega_{L}} \right)^{\frac{\omega_{H}}{\omega_{L} - \omega_{H}}} - \left( \frac{\omega_{H}}{\omega_{L}} \right)^{\frac{\omega_{H}}{\omega_{L} - \omega_{H}}}} \right\rbrack}\end{matrix} & \left( {{Formula}\quad 11} \right)\end{matrix}$

The maximum flow rate produced within the mixer tube, which occurs att=0, is given by; $\begin{matrix}{I_{imp\_ max} = \frac{{- V_{stroke}}{A_{0} \cdot \omega_{n} \cdot \left( {\omega_{H} - \omega_{L}} \right)}}{2 \cdot \sqrt{\zeta^{2} - 1}}} & \left( {{Formula}\quad 12} \right)\end{matrix}$

Control of the performance characteristics of the device can be achievedby careful selection of the parameters of the device. Design inputs,such as, but not limited to, the inner and outer diameters of the tubingin the double-lumen catheter and the cannula interfacing with the bodycavity, the pump frequency and the stroke volume may be set to producethe performance characteristics required for a given design.

Example data for two sets of design inputs, specifically for an examplehigh flow and low flow configuration, can be seen in FIGS. 5A and 5B.The relevant input data and calculations can be found in the spreadsheetof FIGS. 7A-7B. In each case, the stroke volume was set to 0.5 uL. Itcan be seen from the results that the high flow configuration exchangesabout three times the mixer tube volume, while the low flowconfiguration exchanges a volume approximately equal to that of themixer tube. It should be noted that the flow rates vary substantiallywith time. For example, in the high flow configuration, the system draws0.22 uL into the system in approximately 10 sec, but takes approximately1.5 min to fully expel it. By setting the device to operate continuouslyat a predetermined frequency, a relatively small exchange volume (onlyseveral times that of the mixer tube volume) and flow rates is possible.Also, the pump frequency should be slow compared to ω_(L). Thecalculations used in the spreadsheet of FIGS. 7A-7B calculates a pumpfrequency which is 3 times slower than ω_(L). This margin can beadjusted depending on the desired pumping characteristics.

In an alternative embodiment of the invention, the micropump input canbe modulated so that it periodically turns on and off at a frequencymuch lower than the pump cycle frequency, and also more slowly than thelargest system time constant. In this operating mode, the system timeconstants are large relative to both the pulse time and the pump cycleperiod. As a result, the pump effectively looks like a constant current(flow) source rather than a pulse train, as can be seen in FIG. 4B,which depicts a qualitative plot of the pump flow output with respect totime. The resulting flow rate in this configuration is given by;I _(S0) =V _(stroke) ·f _(p)  (Formula 13)where V_(stroke) is the pump's stroke volume and f_(p) the pumpfrequency. In this case, the pump is modeled as a step function current(flow) source, again, assuming it is left “on” longer than the longestsystem time constant.The step input in Laplace domain is given by${I_{S} = \frac{I_{S\quad 0}}{s}},$so Formula (5) becomes; $\begin{matrix}{\frac{I_{US}}{I_{S\quad 0}} = \frac{{- A_{0}} \cdot \omega_{n}^{2}}{\left( {s^{2} + {2 \cdot \zeta \cdot \omega_{n} \cdot s} + \omega_{n}^{2}} \right)}} & \left( {{Formula}\quad 14} \right)\end{matrix}$and the time domain step response is; $\begin{matrix}{I_{us} = {\frac{{- I_{S\quad 0}}{A_{0} \cdot \omega_{n}}}{2 \cdot \sqrt{\zeta^{2} - 1}} \cdot \left( {{\exp\left( {{- \omega_{L}} \cdot t} \right)} - {\exp\left( {{- \omega_{H}} \cdot t} \right)}} \right)}} & \left( {{Formula}\quad 15} \right)\end{matrix}$where I_(S0) is the pump flow rate amplitude. The time dependentresponse approaches zero for large time, due to the decayingexponentials. Its integral, the fluid volume, is given by;$\begin{matrix}{V_{us} = {\frac{{- I_{S\quad 0}} \cdot A_{0} \cdot \omega_{n}}{2 \cdot \sqrt{\zeta^{2} - 1}} \cdot \begin{pmatrix}{\frac{\exp\left( {{- \omega_{H}} \cdot t} \right)}{\omega_{H}} - \frac{\exp\left( {{- \omega_{L}} \cdot t} \right)}{\omega_{L}} +} \\\frac{2 \cdot \sqrt{\zeta^{2} - 1}}{\omega_{n}}\end{pmatrix}}} & \left( {{Formula}\quad 16} \right)\end{matrix}$which asymptotically approaches a constant value, given by;V _(cycus) =−I _(S0) ·A ₀  (Formula 17)

As in the cases with a continuously operating micropump, the maximumflow rate in this configuration is critical to the design, and is givenby; $\begin{matrix}{I_{us\_ max} = {\frac{{- I_{S\quad 0}}{A_{0} \cdot \omega_{n}}}{2 \cdot \sqrt{\zeta^{2} - 1}} \cdot \begin{bmatrix}\left( \frac{\omega_{L}}{\omega_{H}} \right)^{\frac{\omega_{L}}{\omega_{H} - \omega_{L}}} & {- \left( \frac{\omega_{L}}{\omega_{H}} \right)^{\frac{\omega_{H}}{\omega_{H} - \omega_{L}}}}\end{bmatrix}}} & \left( {{Formula}\quad 18} \right)\end{matrix}$

Predicted output data for the configuration where the micropump input ismodulated can be seen in FIGS. 5A and 5B. Again, data is shown for twosets of design inputs, specifically for an example low flow and highflow configuration. The high flow case, shown in FIG. 6B, clearlygenerates a larger volume exchange, but because of the larger systemtime constants, has a lower maximum flow rate than the lower flowconfiguration. In this case, the larger capacitances more thancompensate for the reductions in feed resistances in producing largertime constants and system gain A₀. In using the embodiment wherein themicropump input is modulated, it should be noted that the volumeexchanges shown will be achieved only if the modulation time is largerelative to the slowest time constant.

In choosing system components to optimize performance, it should benoted that the maximum pressure developed at the pump is given by;P _(max) =I _(S0)·(R _(FT) +R _(RT) +R _(F) +R _(R))  (Formula 19)Thus, it is important to choose component values carefully such that thepump will perform properly. It should be noted that In each of the aboveexample embodiments only the primary feed and return lines were assumedto have sufficient compliance to contribute significantly to thecapacitance in the system. The elastic modulus used in the calculationswas 11 MPa.

It can be seen that by modulating the micropump input, substantiallylarger exchange volumes and flow rates are possible than for theexamples where the micropump operates continuously at a predeterminedfrequency. The pump frequency is used to set the average flow rate. Theexchange volume and flow rate are directly proportional to this averageflow rate and thus the pump frequency. It should be noted that theequations and example data shown here are only accurate for caseswherein the pump frequency is not comparable to ω_(L). Specifically, theerror is 10% at f_(P)=2ω_(L). and will increase as f_(P) is decreased.Further, the pump on-off modulation frequency should be slow compared toω_(L). The example data is calculated for a pump modulation frequencywhich is three times slower than ω_(L). This margin can be adjusteddepending on the desired pumping characteristics.

A possible disadvantage of the embodiments of the drug deliveryapparatus described above with reference to FIGS. 2A-2B and 3A-3C is itssensitivity to tubing dimensions, which may change over time.Additionally, a micropump employed by the apparatus according to thoseembodiments should be biocompatible and suitable for pumping biologicalfluids. Accordingly, in alternative embodiments, instead of acirculating fluid loop, a flexible diaphragm or slidable piston incontact with a working chamber facilitates reciprocating flow throughthe cannula and into and out of the body organ, simplifyingimplementation of the drug delivery apparatus as a microfabricatedsystem and improving its resistance to biological fouling.

In particular, referring to FIGS. 8A-8B, in some embodiments, a drugdelivery apparatus 800 includes a variable-volume vessel 805 and aninterface member 810, for example, a cannula. The apparatus may alsoinclude a reservoir member 815 for storing a drug 817. The vessel has aworking chamber 820 in fluid communication with the interface member andthe reservoir member, if one is used by the apparatus. The reservoirmember is separated from the chamber by a valve 825. The working chambercontains a therapeutic fluid 830, including the drug and, in operation,certain amount of bodily fluid, for example, perilymph. The apparatusfurther includes an actuator 835 for varying the pressure within theworking chamber by altering its volume. For example, in many versions ofthese embodiments, the actuator periodically increases and decreases thevolume of the chamber by either deflecting or slidably moving at least aportion of at least one wall of the vessel 805. In some versions, eitheran entire wall 840 of the vessel 805, or a portion thereof, includes, isa flexible membrane or diaphragm, as shown in FIGS. 8A-8B. In otherversions of these embodiments, the vessel 805 has a slidably movablewall. The motion of the wall 840 could be produced by a mechanicalactuator, such as a miniature electromagnetic actuator disclosed in aco-pending patent application Ser. No. 11/169,211 entitled“Electromagnetically-Actuated Microfluidic Flow Regulators and RelatedApplications” and incorporated herein by reference. Alternatively, thewall motion could be produced by fluid pressure from a conventionalpneumatic or hydraulic apparatus, for example as shown in FIG. 8B,employing a working fluid or gas 845, a pump 850, and a valve 855. Otherconfigurations of the vessel 805 facilitating varying the pressurewithin the working chamber by altering its volume are also contemplated.

In operation, similarly to the embodiments described above withreference to FIGS. 2A-2B and 3A-3C, the interface member 810, forexample, a cannula, is brought into contact with a desired bodily cavityof a patient, for example, a cochlea of a human ear, in order toperiodically deliver the therapeutic fluid to and draw bodily fluid fromthe bodily cavity. At desired intervals, the actuator 835 causes thewall 840 or a portion thereof to move or deflect inward, reducing thevolume of the chamber 820 and causing therapeutic fluid 830 to flow fromthe chamber 820 through the cannula to the patient's bodily cavity. Atopposing intervals, the the wall 840 or a portion thereof moves ordeflects in the opposite direction, causing bodily fluid to flow fromthe patient's bodily cavity through the cannula to the chamber. Thus,composition of the therapeutic fluid, including concentration of thedrug therein, varies during its recirculation through the workingchamber. At other intervals, which may or may not be synchronized withor coupled to the periodic motion of the wall 840, the reservoir 815releases drug compounds to be mixed with the bodily fluid in thechamber. Alternatively, pressure variation in the working chamber due tothe motion of wall 840 could be used in a mechanism for dispensing drugfrom the reservoir. For example, synchronized operation of the reservoirvalve with the low-pressure phase in the chamber could cause dispensingof drug.

Similarly to the embodiments described above with reference to FIGS.2A-2B and 3A-3C, the apparatus may also include a regulating systemand/or a control system in communication with the vessel 805, theactuator 835 and/or the reservoir 815 for monitoring and maintaining adesirable drug delivery rate and/or controlling a flow pattern of thefluids through the working chamber. Further, the variable-volume vesseland the actuator can be shaped and dimensioned to fit within a desiredbodily cavity, for example, a mastoid cavity of a human.

The invention may be embodied in other specific forms without departingform the spirit or essential characteristics thereof The foregoingembodiments, therefore, are to be considered in all respectsillustrative rather than limiting on the invention described herein.Scope of the invention is thus indicated by the appended claims ratherthan by the foregoing description, and all changes that come within themeaning and range of equivalency of the claims are intended to beembraced therein.

1. A drug delivery apparatus for delivering a drug into a bodily fluidin a bodily cavity of a patient over a period of time, the apparatuscomprising: a variable-volume vessel defining a working chamber forreceiving a drug and recirculating a therapeutic fluid, the therapeuticfluid comprising a first volume of the bodily fluid and the drugcontained therein; an actuator for varying the pressure within thechamber by altering the volume of the chamber; and an interface memberin fluid communication with the chamber and the bodily cavity forperiodically delivering at least a portion of the therapeutic fluid to apredetermined location in the bodily cavity and drawing a second volumeof the bodily fluid from the bodily cavity.
 2. The apparatus of claim 1wherein the interface member comprises a cannula.
 3. The apparatus ofclaim 1 wherein the interface member is in fluid communication with acochlea of a human ear.
 4. The apparatus of claim 1 whereinconcentration of the drug in the therapeutic fluid varies duringrecirculation thereof through the working chamber.
 5. The apparatus ofclaim 1, further comprising a reservoir member for storing the drug, thereservoir member defining a storage chamber in communication with theworking chamber.
 6. The apparatus of claim 1 wherein the actuatorperiodically increases and decreases the volume of the chamber by atleast one of deflecting or slidably moving at least a portion of atleast one wall of the variable-volume vessel.
 7. The apparatus of claim1 wherein the variable-volume vessel comprises a slidably movable wall.8. The apparatus of claim 1 wherein the variable-volume vessel comprisesa wall having a deflectable portion.
 9. The apparatus of claim 8 whereinthe wall comprises a flexible membrane.
 10. The apparatus of claim 1wherein the actuator is selected from the group consisting of:electromagnetic, pneumatic, or hydraulic devices
 11. The apparatus ofclaim 1 wherein the actuator causes the therapeutic fluid to flowthrough the working chamber at the flow rate of less than about onemicroliter per minute.
 12. The apparatus of claim 1, further comprisinga regulating system in communication at least with the actuator formaintaining a desirable drug delivery rate.
 13. The apparatus of claim12 wherein the regulating system comprises a sensor for periodicallymeasuring concentration of drug in the therapeutic fluid andtransmitting the measured value of the concentration to the regulatingsystem.
 14. The apparatus of claim 12 wherein the regulating systemcomprises a sensor for periodically measuring concentration of drug inthe bodily fluid and transmitting the measured value of theconcentration to the regulating system.
 15. The apparatus of claim 1,further comprising a control system in electric communication with theactuator for controlling a flow pattern of the therapeutic fluid throughthe working chamber.
 16. The apparatus of claim 1 wherein thevariable-volume vessel and the actuator are shaped and dimensioned tofit within a mastoid cavity of a human.
 17. A method for delivering adrug into a bodily fluid in a bodily cavity of a patient over a periodof time, the method comprising the steps of: (a) drawing a first volumeof the bodily fluid from the patient's bodily cavity through aninterface member directly into a variable-volume vessel defining aworking chamber; (b) mixing the drug with the bodily fluid in theworking chamber thereby obtaining a therapeutic fluid; and (c)increasing a pressure within the working chamber to release a firstvolume of the therapeutic fluid into the bodily cavity.
 18. The methodof claim 17 wherein the bodily cavity is a human cochlea and the bodilyfluid comprises human perilymph.
 19. The method of claim 17, furthercomprising controllably recirculating the therapeutic fluid through theinterface member between the working chamber and the bodily cavity. 20.The method of claim 19, further comprising altering concentration of thedrug in the therapeutic fluid during recirculation thereof.